Recently, the implantation of sophisticated electronic devices has become a routine procedure in many therapeutic procedures. Today cerebellar, bladder, peroneal nerve, and spinal cord stimulator, as well as cardiac pacemaker, implants are commercially available. Cochlear implants and other devices are in the development stage and will be useful treatment devices in the future.
Although these devices are used in various organs of the body, the applications have many elements in common. Each implantable device requires a signal generating system and a signal delivery system. In order to design a reliable, safe therapeutic device with a long life span for any of these applications, an understanding of the electrode-tissue interface is essential.
The main factors to be considered with regard to implantable electrodes are their constituent material, size and shape, the electrolyte composition, the system electrochemistry, and the tissue reaction. Each design factor must be tailored to meet the requirements of the application. For example, when implanting a cardiac pacemaker electrode, a degree of imprecision in the location of the electrode is tolerable, whereas difficulties in fixation of the implant are extremely undesirable. With cochlear implants, the electrode is easily fixed into place but the range of tolerance for positioning is extremely narrow.
The electronic design of an implantable stimulation apparatus is influenced by the optimum signal for tissue stimulation. In the case of a cardiac pacemaker, implant lifetime is determined by the energy delivered per pulse, and the pacemaker will have a longer life if the energy delivered per pulse is maintained at a minimum.
In physiological terms, the cardiac pacemaker must be capable of generating a signal with a sufficient magnitude to depolarize the excitable cells. The electrode size and shape, the electrolyte conductivity, and the distance separating the electrode and the excitable tissue determine the energy required of the pacemaker.
Over the last few years platinum and platinum in combination with iridium have become the most widely used cardiac pacemaker electrode materials, and have been accepted by regulatory bodies. Gold, stainless steel, palladium, silver, titanium, carbon, and tantalum have also been used in experimental applications.
In order to evaluate the suitability of an electrode system for each application, the impedance of the system to signals of differing frequencies must be known. In the case of a pacemaker, the current drain, and therefore the implant lifetime, is determined by the impedance to pacing pulses. The pacemaker electrode must not only deliver a pacing pulse with a pulse width in the range of 0.1-2.0 msec. to the tissue, but must also transmit a QRS signal (50 Hz) to the pacer circuitry. The pacemaker pulse is inhibited when normal ventricular depolarizations occur. The electrode-electrolyte system impedance is higher for sensing than for pacing. Electrodes are also used for pacing and sensing in the atrium which exhibits different stimulation and depolarization parameters than those of the ventricle.
The electrode/tissue system impedance characteristics may be understood in terms of an interface component which is the dominant component and occurs within one micron of the surface of the electrode, and a spreading resistance which depends predominantly on the tissue resistivity. The former reflects the charge transfer characteristics of the interface, and the latter reflects the size and shape of the electrode and the resistivity of the tissue.
The magnitude of the electrode/electrolyte impedance is frequency dependent because of the polarization effects at the interface with the surrounding tissue. At low frequencies, the interface impedance is significant.
The current drain of a pacemaker is determined by the impedance of the pacemaker circuitry, the nature of the electrode lead resistance, and the characteristics of the electrode tip interface with the electrolyte system. Since for a given pacemaker circuit and electrode lead design the current drain is well defined, the nature of the electrode tip/tissue interface determines the overall current requirements of the system. The most significant frequency of the pacing pulse is in the order of 1 KHz. At this frequency, the interface impedance is small and most of the impedance to the pacing pulses is due to the bulk or spreading impedance. This is determined by the shape of the electrode tip and is inversely related to the radius of the electrode tip.
The pacing impedance is indicative of the geometric surface area of the electrode tip and is a function of the electrode radius. For example, a hemispherical electrode tip having a small radius will have a higher pacing impedance and smaller current drain than a similarly shaped electrode tip of larger radius.
The most significant frequency components of a signal to be sensed, i.e., the ventricular QRS, are in the bandwidth of 20-100 Hz. In this region, the interface impedance becomes most significant. The interface impedance is determined by the microsurface area of the electrode tip and develops within a few microns of the surface. The microsurface area of an electrode tip is the area which includes all of the ridges, crevices, and indentations in the surface of the electrode tip.
It has been determined that the pacing threshold is a reflection of the energy required for a pulse to initiate a cardiac contraction. The stimulation threshold rises for weeks after the implant of a cardiac pacemaker as a result of an increase in the spacing between the electrode and the excitable tissue. The increase occurs due to the development of a fibrous capsule around the electrode tip which is reported to be between 0.3 mm and 3 mm thick.
In view of the above characteristics of an electrode for a cardiac pacemaker, it is clear that an electrode tip with a small geometric surface area and high pacing impedance will have a low current drain. However, in order to enhance sensing, the same electrode tip should have a large microsurface area to result in a low sensing impedance. Although this combination of characteristics seems to be incompatible, it has been obtained in a cardiac pacemaker electrode tip that is constructed to be porous. One such porous electrode tip comprises a platinum-iridium screen covering a mesh ball of the same alloy.
One of the advantages of a porous electrode tip is the ability to minimize the radius of the electrode tip to present a small geometric surface area while having an increased aggregate microsurface area. This extends the lifetime of the pacemaker by providing a high pacing impedance and a lower current drain. On the other hand, a large microsurface area is provided by virtue of the mesh-like construction. The large microsurface area enables enhanced sensing by lowering the sensing impedance.
It has been found, however, that a mesh-like construction presents difficulties in fabricating the porous electrode tip. These include the control over the shaping and the bonding of the mesh in a precise manner. These attributes are necessary to make certain that the porous electrode tip will exhibit uniform impedance and other electrical properties as well as sufficient reliability of construction.
Thus, there has been a need to provide a cardiac pacemaker porous electrode tip which is easier to make in appropriate shapes, has predictable, predetermined electrical properties, and is of adequate reliability.
An area of vital importance that is often not considered in detail is the tissue reaction to the electrode itself. Biocompatibility is a term often used to describe the general suitability of a material for implantation. Generalizations concerning this term are of little practical value because each implant application has different tissue reaction requirements. Systemic toxicity is of course undesirable, however, fibrous reaction may be useful for prosthesis attachment. In the case of a pacemaker electrode, minimal tissue reaction is desired around the tip but firm attachment of the electrode to the tissue is essential.
A porous electrode tip allows rapid fibrous tissue growth into a hollow area or cavity in the electrode tip itself to enhance attachment of the electrode to the heart. A smaller dislodgement rate is expected as a result of such tissue ingrowth.
A further aspect of importance is selection of pore size, which must be such as to accommodate economical construction techniques, overall dimensional tolerances, and tissue response constraints. Any design must balance the pore size, the number of pores, the pore intercommunications, and mechanical stability. Recent research has suggested that the selection of pore size has an influence on the tissue reaction to porous materials. Although this area is in its infancy, it is clear that pore diameter should be at least 15 microns to allow tissue ingrowth. It has also been suggested that pore size will determine tissue capsule thickness and, thus, stimulation threshold.
From the above discussion, it is clear that an ideal porous electrode tip construction should allow precise selection of pore size and pore configuration. The invention to be described provides a clear improvement over the prior art in this regard.
The present invention allows the independent variation of porosity and microsurface area. The pore size may be selected on the basis of optimal tissue ingrowth and attachment with minimal chronic tissue thickness. The high microsurface area decreases the sensing impedance and is determined predominantly by the roughening process. Theoretical studies and laboratory experiments have shown that with the pore size selected, the internal cap surface, which defines an internal cavity, does not make a major contribution to the reduction of sensing impedance.
The importance of distinguishing between calculated and available microsurface area for sensing is not stated in the literature relating to the prior art. The embodiment to be described hereinafter takes advantage of the microsurface area of a roughened external surface as well as the surface area of the cap. The roughening process is critical to the achievement of sufficiently low sensing impendance and this development gives rise to improved porous electrode design and construction possibilities.
Additional objects and advantages of the invention will be set forth in part in the description which follows, and in part will be apparent from the description or may be learned from the practice of the invention. The objects and advantages of the invention may be realized and obtained by means of the instrumentalities and combinations particularly pointed out in the appended claims.